Field of the Invention
The present invention relates to x-ray and gamma-ray generation via laser Compton scattering and more specifically, it relates to subtraction radiology utilizing laser-Compton x-ray sources.
Description of Related Art
In conventional 2-D x-ray/gamma-ray imaging, the patient or object is illuminated with a flat field of x-rays or gamma-rays and the transmitted signal is recorded on a 2D film or array of detectors. Variations of material density within the object cause variations in beam transmission for the penetrating radiation and these variations appear as shadows on film or a detector array. The dynamic range of this imaging technique is determined by the response function of the detector system and by the object thickness and secondary x-ray scattering by the object. In addition, all parts of the object see the same input flux (photons per unit area) and the total dose impinging upon the object is set by the area of the object and by the flux required to penetrate the densest region of the object, i.e., the flux required to resolve the structures of interest within the object. In this imaging modality, the entire object sees a high dose.
For some imaging procedures in which the desired object is either small or low density, a higher atomic number contrast agent is injected or ingested to provide specific information about targeted structures. For example in coronary angiography the goal is to image the blood vessels and in particular to locate areas of reduced blood vessel aperture or blockages. Because the blood and the blood vessels are soft tissue and small in size, the total x-ray attenuation by them is small compared to the background matrix in which they are present and thus it is hard if not impossible to sufficiently resolve them in a conventional, whole body x-ray image. To overcome this issue, a dense material generally of higher atomic number than the surrounding biological material is injected into the blood stream to increase the x-ray attenuation in the areas of interest and in doing so improve contrast. Contrast agents used in human imaging tasks must of course be certified as being biologically inert or at least relatively so. For coronary angiography, iodine-containing compounds have been used as contrast agents. It should be noted that while this procedure does improve contrast and provide the required spatial information, the dose received by the patient can be very high. Some coronary angiography procedures can expose the patient to a full year's allowable dose.
In order to increase the contrast and/or reduce the required dose for an image at a desired contrast level, two-color subtraction imaging has been suggested and demonstrated. In this modality, the patient is illuminated twice with a tunable, quasi-mono-energetic x-ray source. In one case the x-ray source has its energy set slightly above the k-shell absorption edge of the contrast agent and in the other case it is set slightly below. As shown in FIG. 1, the absorption cross section for the contrast material varies dramatically around the k-shell absorption region while the absorption cross section for the surrounding material can be relatively unchanged. If the two images are normalized to have the same signal in regions not containing the contrast agent, then a subtraction of the normalized images will be an image whose content is due primarily to the contrast agent.
While early experiments conducted with filtered light from synchrotron x-ray sources demonstrated that this procedure could dramatically increase image contrast and/or reduce dose to the patient, its implementation in real-world clinical environments has been relatively limited due to the lack of clinically-compatible, quasi-mono-energetic x-ray sources. Synchrotron sources are expensive (>$100M), large (>100 m in diameter) and relatively uncommon. In addition the output from a synchrotron source is constant and not rapidly adjustable nor easily scanned across the object.
It should also be noted that some have attempted to use conventional bremsstrahlung sources for k-edge imaging by changing the end point energy of the electron beam impinging upon the rotating anode so that the highest energy photons are either above or slightly below the desired k-edge absorption. In practice this, however, does not work very well as the total x-ray content of a bremsstrahlung source extends from the end point energy of the electron beam to DC, thus the fraction of the beam spectrum that is above the k-edge is relatively small compared to the total x-ray production and the image is thus dominated by background absorption. The dose to the patient is also high in this mode as it predominantly comes from the low energy tail of the bremsstrahlung spectrum of the source. To some extent this issue can be minimized by attenuating the beam with a low atomic number material that preferentially reduces the low energy portion of the spectrum relative to the high-energy portion but this of course reduces the total x-ray flux available for imaging, increases the proportion of image-degrading, scattered x-ray content within the illuminating x-ray beam and requires a higher current anode device to create the same number of useful above and below k-edge photons at the object.
Note that the k-shell edge and not the outer shell absorption edges, i.e. L and M is generally used for two color clinical imaging as the x-ray energy required to remove a k-shell electron generally falls in the x-ray region of interest to clinical radiography while the outer shell absorptions occur at lower x-ray energies. The same two-color image subtraction scheme can be implemented, however, at lower energies using outer shell absorption edges if object and source are compatible.